The present invention is generally related to an X-ray tube assembly and more particularly to an X-ray tube assembly in which anode or cathode elements are moveable with respect to each other.
Radiographic imaging systems, such as X-ray and computed tomography (CT) have been employed for observing interior aspects of an object. Typically, the imaging systems include an X-ray source that is configured to emit X-rays toward an object of interest, such as a patient. A detecting device, such as an array of radiation detectors, is positioned on the other side of the object and is configured to detect the X-rays transmitted through the object of interest.
X-ray or radiographic imaging is the basis of a number of diagnostic imaging systems. Computed tomography (CT) is one example of such a system that is predicated upon the acquisition of data using the principles of radiography. By mapping the incomplete penetration of X-rays through an object from multiple different angles, CT image systems generate volumetric representations of the object. Dedicated Breast Tomography (DBT) operates under similar principles to CT, but, whereas CT generally acquires data a full 360 degree range of angles about the imaged object, DBT will acquire only a limited range of angles such as 30 to 120 degrees and with a system geometry optimized to breast imaging, Typically, in CT and DBT imaging systems, a single X-ray source emits a single fan-shaped or cone-shaped beam toward a subject or object, such as a patient or a piece of luggage. The imaging angle is defined relative to the fixed imaged object by a line that joins the focal spot location in the tube to the detector center. In a moving source configuration, the angle is changed by moving the source relative to the fixed subject and collecting data at times. In a distributed source configuration, multiple focal spot locations are provided in the system and are energized in a sequence so as to provide data at different imaging angles. The beam, after being attenuated by the subject, impinges upon an array of radiation elements. One embodiment of this array could be a pixel array within a radiographic detector or multiple pixels spread over multiple discreet detectors. The intensity of the attenuated beam radiation received at the detector array is typically dependent upon the attenuation of the X-ray beam by the subject. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately produces an image.
Generally, for CT the X-ray source and the detector array are rotated about an axis within an imaging plane and around the subject. For DBT, the X-ray source is moved relative to the subject and the detector may or may not move. X-ray sources typically include X-ray tubes, which emit the X-ray beam at a focal point. X-ray detectors typically include a collimator for collimating X-ray beams received at the detector, a scintillator for converting X-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator and producing electrical signals therefrom. Alternately, a photoconductor layer can serve a combined function of the scintillator and photodiodes by directly absorbing X-rays and converting directly to electrical signals.
In an indirect detector type, the scintillator converts X-rays to light energy. The scintillator discharges light energy to a photodiode adjacent thereto. Each photodiode detects the light energy and generates a corresponding electrical signal. In a direct detector type, the photoconductor converts X-rays to electric charge in proportion to the energy and delivers that charge under the influence of a high voltage to electrical contacts. The outputs of the photodiodes or photoconductor are then transmitted to the data processing system for image reconstruction. Other types of radiation detectors exist such as those that convert X-rays directly to electrical signals without the use of a scintillator or photodiode.
CT systems, as well as X-ray systems, typically utilize a rotating anode during the data acquisition process. Rotating the anode presents a moving electrode track at the spot which generates the X-ray fan beam, and thereby spreads the thermal load on an enlarged surface area of the anode. That is, the anode typically includes a single target electrode that is mounted or integrated with an anode disc. The anode disc may be rotated by an induction motor during data acquisition. Rotation is initiated before the start of a mechanical scan of the source. The rotation rate is maintained at a constant rate throughout the source trajectory and not coordinated specifically with the acquisitions of data at multiple locations. Since the electrons striking the anode deposit most of their energy as heat, with a small fraction emitted as X-rays, producing X-rays in quantities sufficient for acceptable image quality generates a large amount of heat. It can be the case that the heat is generated sufficiently rapidly and in a small focal area on the anode electrode surface as to raise the temperature close to the melting temperature of the electrode material or creating destructive thermal strain and shock. The consequences of such thermal overloading can be melting or cracking on the anode surface. A number of techniques have been developed to accommodate the thermal load placed on the anode during the X-ray generate process.
For example, advancements in the detection of X-ray attenuation has allowed for a reduction in X-ray dose necessary for delivering images with sufficient quality. X-ray dose delivered to the patient and tube current are directly related and, as such, a reduction in tube current results in a reduction in X-ray dosage. Similarly, a drop in tube current, i.e. reduction in the number of striking electrons on the anode target, reduces the thermal load placed on the anode target during data acquisition. Simply, less power is needed to generate the X-ray necessary for data acquisition. However, counteracting the advances in detection equipment is the need to decrease scan time or increase throughput (e.g. patients per hour) that call for increased X-ray power. X-rays are generated at a focal local on the anode electrode as a result of electrons emitted from a cathode striking a target electrode mounted to or integrated with the anode disc. A control circuit sets the number of electrons emitted from the cathode and the voltage potential placed across the cathode and anode in order meet the imaging requirements and to protect the tube from over-heating. Total heat load on the anode that is proportional to the emitted current and the voltage potential placed across the cathode and anode. Optimal patient safety conditions can call for higher tube voltage or pre-patient attenuation filters that drive higher X-ray power requirements that challenge the thermal condition of the anode. In spite of advances in detection equipment, a minimum number of electrons must be generated for meaningful data acquisition under the constraints of minimum patient dose or increased throughput. As a result, a mere reduction in tube current is insufficient to address the thermal load on the anode resulting from X-ray generation.
Another approach is predicated upon the spreading of the generated heat across the surface and mass of the anode disc. By rotating the anode disc as electrons are striking the target electrode, the heat generated therefrom may be spread across a track of the anode disc rather than at a single focal location on the target electrode surface. This rotation of the anode disc effectively reduces the thermal load placed at any single location on target electrode. As a result, tube current may be increased without thermal overloading of the anode. Generally, the faster the anode disc is rotated the higher the tube current that may be used.
Increasing the tube current and effectively the power levels of the X-ray tube assembly is particularly desirable for short duration high power reconstruction protocols. With these protocols, the gantry is caused to rotate at significantly fast rotational speeds. Through increased rotational gantry speed, the overall exam time may be decreased. Decreasing the overall exam or scan time improves patient throughput and reduces patient discomfort which reduces patient-induced motion artifacts in the reconstructed image. To support faster gantry speeds, the X-ray tube must output sufficiently more instantaneous power which is required for short duration protocols.
To provide the requisite instantaneous power needed for short duration protocols, the X-ray tube must output more power without exceeding the thermal load of the target electrode. As mentioned above, rotating the anode disc during X-ray generation reduces the thermal load on the electrode target.